Nuclear medical diagnosis apparatus

ABSTRACT

A PET apparatus comprises a plurality of detector units in the circumferential direction, wherein the detector unit includes a plurality of unit substrates therein, and wherein the unit substrate includes: a plurality of detectors upon which a γ-ray is incident; and an analog ASIC and digital ASIC for processing a γ-ray detection signal outputted by each of the detectors. The analog ASIC includes two slow systems having mutually different time constants, each of which outputs a pulseheight value. A noise determination part of the digital ASIC determines whether a relevant detection signal is an intended γ-ray detection signal or a noise based on a correlation between the pulseheight values, and a noise counting part counts the number of times of noise determination, and a detector output signal processing control part controls the signal processing with respect to an output signal from a relevant detector based on the count.

BACKGROUND OF THE INVENTION

The present invention relates to nuclear medical diagnosis apparatuses,and in particular relates to a positron emission computed tomographyapparatus (hereinafter, referred to as a PET apparatus) that is one typeof the nuclear medical diagnosis apparatuses using radiation detectors,a single photon emission computed tomography apparatus (hereinafter,referred to as a SPECT apparatus), and a γ-camera using γ-rays passingthrough a test object.

Conventionally, as the radiation detectors for detecting radiation, suchas γ-rays, a radiation detector using an NaI scintillator is known. Inthe γ-camera equipped with an NaI scintillator, radiation (γ-ray) isincident upon the NaI scintillator at an angle limited by a large numberof collimators, thus interacting with an NaI crystal to emitscintillation light. This light reaches a photomultiplier or aphotodiode via a light guide to become an electrical signal. Theelectrical signal is shaped by a measuring circuit mounted on a circuitsubstrate and is then sent to an external data acquisition system froman output connector. In addition, these NaI scintillator, light guide,photomultiplier, measuring circuit, circuit substrate, and the like areentirely housed in a light shielding case to block electromagnetic wavesother than external γ-rays.

In addition, here, when an NaI scintillator is combined with aphotomultiplier or a photodiode, this combination is defined as aradiation detector.

The radiation detector that detects radiation based on a principledifferent from the principle of the radiation detector combining suchNaI scintillator with a photomultiplier or a photodiode is asemiconductor radiation detector equipped with semiconductor radiationdetection elements using a semiconductor material, such as CdTe (cadmiumtelluride), CdZnTe (zinc telluride cadmium), HgI₂ (mercury iodide), TlBr(thallium bromide), or GaAs (gallium arsenide). In this semiconductorradiation detector, the semiconductor radiation detection elementconvert a charge resulting from an interaction between a γ-ray and thesemiconductor material into an electrical signal, so this semiconductorradiation detector can accomplish conversions into an electric signalmore efficiently than the scintillator can, and also accomplishminiaturization. Accordingly, this semiconductor radiation detectorattracts much attention.

In addition, in the nuclear medical diagnosis apparatus that generatesan image using a large number of such radiation detectors, there is aproblem that when a noise signal mixing with an intended radiationdetection signal outputted from the radiation detector is outputted froman abnormal radiation detector, i.e., a photomultiplier, a photodiode,or a semiconductor radiation detector, this noise signal is alsoprocessed during image generation.

For this reason, the conventional technique described in JP-A-2006-98411(paragraphs [0032], and to [0039]) uses the method for detecting anabnormal semiconductor radiation detector and excluding an output signaltherefrom, wherein in a PET apparatus or SPECT apparatus equipped with aplurality of semiconductor radiation detectors, semiconductor radiationdetectors are arranged so as to surround the circumference of the bodyaxis of a test object and also to be in multiple layers in the radialdirection. Here, with respect to γ-rays in the radial direction passingthrough the semiconductor radiation detectors in multiple layers, aratio between signals outputted by the semiconductor radiation detectorsin each layer is used to determine which semiconductor radiationdetector on which layer is abnormal when the ratio deviates from apredetermined ratio by a specified amount or more.

In addition, in the technique described in the above-describedJP-A-2006-98411, an abnormal radiation detector can be determined onlyafter carrying out transmission imaging or actual SPECT imaging, or PETimaging, or imaging of transmission image, and even if a noise signal isoutputted from an abnormal radiation detector, a signal processingdevice will process an output signal including this noise signal as aradiation detection signal, thus resulting in an increase in the signalprocessing load of the signal processing device. Moreover, when multiplelayers of semiconductor radiation detectors are disposed in the radialdirection, multiple layers of semiconductor radiation detectors in theradial direction at a specific position can be determined as abnormal,but this technique can not be applied to the case where radiationdetectors using a scintillator are arranged in one layer in the radialdirection.

As a result, there is a problem that the generation of a SPECT image,PET image, and transmission image can not be started until thedetermination of a abnormal radiation detector is completed in dataprocessing after the imaging.

SUMMARY OF THE INVENTION

The present invention has been made in view of the above problem, and itis an object of the present invention to provide a nuclear medicaldiagnosis apparatus that can, even without disposing multiple layers ofradiation detectors, determine an abnormal radiation detector duringimage pickup and eliminate an adverse effect on the generated image dueto a noise signal from the radiation detector.

In order to achieve the above-described objects, according to an aspectof the present invention, a signal processing device comprises: adetermination unit that determines whether an output signal from aradiation detector is an intended radiation detection signal or a noise;a counting unit that counts for each radiation detector the number oftimes the output signal from the detector is determined as a noise; anda control unit which, based on the number of times the output signalfrom the detector is determined as a noise, determines the relevantradiation detector as faulty and controls so as not to process an outputsignal from the radiation detector that is determined as faulty.

The present invention can provide a nuclear medical diagnosis apparatusthat can determine a abnormal radiation detector during image pickup andeliminate an adverse effect on the generated image due to a noise signalfrom the radiation detector.

Other objects, features and advantages of the invention will becomeapparent from the following description of the embodiments of theinvention taken in conjunction with the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view showing the configuration of a PETapparatus as a nuclear medical diagnosis apparatus concerning a firstembodiment of the present invention.

FIG. 2 is a view schematically showing a cross section in thecircumferential direction of a camera of the PET apparatus of FIG. 1.

FIG. 3A is a partially cutaway perspective view of the camera, showing aconfiguration to mount a detector unit into the camera, and FIG. 3B is across sectional view showing the mounting state of the detector units ina central portion of the camera.

FIG. 4 is a view schematically showing the configuration of asemiconductor radiation detector.

FIG. 5 shows a perspective view in (a) schematically showing constituentelements of the semiconductor radiation detector in a laminatedstructure of FIG. 4, and shows a perspective view of the laminatedstructure made by combining the constituent elements in (b).

FIG. 6 shows a front view in (a) showing a coupling substrate made bycoupling a detector substrate of the semiconductor radiation detectorconcerning the first embodiment with an ASIC substrate, and shows theside view of FIG. 6 (a) in (b).

FIG. 7 is a block diagram showing a schematic configuration of thedigital ASIC, and a connection relation between an analog ASIC and thedigital ASIC.

FIG. 8 is a block diagram schematically showing an analog signalprocessing circuit of the analog ASIC.

FIG. 9 is a block diagram schematically showing the detection signalprocessing and the like of the digital ASIC.

FIG. 10 is a transparent perspective view showing the structure of adetector unit housing a plurality of coupling substrates therein.

FIG. 11 is a side view looking at the detector unit of FIG. 10 in thebody axis direction.

FIG. 12A and FIG. 12B describe that the output timing of a timing signalVT differs depending on positions between an anode and a cathode of adetector that interacts with γ-rays, and FIG. 12A is a detail view of acircuit of a portion outputting the timing signal VT, wherein thesemiconductor radiation detector of FIG. 4 and FIG. 5 is schematicallyshown in a single layer structure, and FIG. 12B is a view showing arelation among the waveform of an output current pulse IA of thedetector, the waveform of an output voltage VB of a preamplifier, andthe waveform of the output voltage VT of a comparator.

FIG. 13A to FIG. 13C describe that the output timing of the timingsignal VT differs depending on a γ-ray detection energy which a detectordetects, wherein FIG. 13A is a spectrum showing the detection energy ofthe detector and the detection frequency, FIG. 13B is a detail view of acircuit of a portion outputting the timing signal VT, wherein thesemiconductor radiation detector of FIG. 4 and FIG. 5 is schematicallyshown in a single layer structure, and FIG. 13C is a view showing arelation among the waveform of the output current pulse IA of thedetector, the waveform of the output voltage VB of the preamplifier, andthe waveform of the output voltage VT of the comparator, and inparticular is a view showing that the rising of the waveform of theoutput current pulse IA differs when the detection energy differs inspite of the same contribution of electron.

FIG. 14 is a view showing the output signal VB of the preamplifier ofFIG. 8, and the output signals VS1 and VE1 of a waveform shaper circuitand a peak hold circuit in a first slow system, respectively, and theoutput signals VS2 and VE2 of the waveform shaper circuit and the peakhold circuit in a second slow system, respectively.

FIG. 15 is a correlation between the pulseheight values VE1, VE2outputted from the first slow system and the second slow system, thecorrelation being used in discrimination between a γ-ray detectionsignal and a noise in a noise determination part in the firstembodiment.

FIG. 16 is a correction value table based on the pulseheight values VE1,VE2 outputted from the first slow system and the second slow system, thepulseheight values being used for correction of a detection time in adetection time correcting part in the first embodiment.

FIG. 17 is a perspective view showing the configuration of a SPECTapparatus as a nuclear medical diagnosis apparatus concerning a secondembodiment of the present invention.

FIG. 18 is a block diagram showing a schematic configuration of adigital ASIC in the SPECT apparatus of FIG. 17, and a connectionrelation between an analog ASIC and the digital ASIC.

FIG. 19 is a block diagram schematically showing a circuit configurationof the analog ASIC in the SPECT apparatus of FIG. 17.

FIG. 20 is a block diagram schematically showing a circuit configurationof the digital ASIC in the SPECT apparatus of FIG. 17.

FIG. 21 shows a correlation between the pulseheight values VE1, VE2outputted from the first slow system and the second slow system, thecorrelation being used in discrimination between a γ-ray detectionsignal and a noise in a noise determination part, when a combination ofa scintillator, and a photomultiplier or a photodiode is used as the γdetector.

DESCRIPTION OF THE INVENTION First Embodiment

Next, a nuclear medical diagnosis apparatus, which is a suitableembodiment of the present invention, will be described suitablyreferring to the accompanying drawings.

In the followings, a nuclear medical diagnosis apparatus of the presentembodiment, elements applied to the present embodiment, such asarrangement (layout) of each device such as analog ASIC onto asubstrate, unitization of substrates, a method for determining a noise,and a method for controlling a radiation detector that is determined asfaulty will be described.

In addition, the analog ASIC means an ASIC (Application SpecificIntegrated Circuit), i.e., an IC for a specific application, forprocessing analog signals, and is one type of LSI (Large ScaleIntegrated Circuit).

(Nuclear Medical Diagnosis Apparatus)

First, the nuclear medical diagnosis apparatus of the present embodimentis described.

As shown in FIG. 1, a PET apparatus 1 as the nuclear medical diagnosisapparatus comprises a camera (image pickup device) 11, a data processingdevice 12, and an operator console 13, and the like. A test object P(see FIG. 2) is placed on a bed 14 and is imaged with the camera 11.

The operator console 13 comprises a display device 13 a that displaystomographic images of the PET apparatus 1, status check results of thePET apparatus 1, and the like, and an input operation part (input unit)13 b, such as a keyboard and a mouse.

As shown in FIG. 2, inside the camera 11, in order to detect γ-rays(radiation) emitted from the test object P, a large number of detectorunits 2, each housing a plurality of coupling substrates (unitsubstrates) 20 (see FIG. 6 for details) provided with a large number ofsemiconductor radiation detectors (radiation detectors, hereinafter,simply referred to as detectors) 21 (see FIG. 4, FIG. 5), are arrangedin the circumferential direction, whereby the γ-rays emitted from thebody of the test object P are detected with the detectors 21.

The detector unit 2 includes, on the coupling substrate 20, anintegrated circuit (ASIC) for measuring the detection energy anddetection time of a γ-ray. Here, the detector unit 2 measures thedetection energy and detection time of a detected γ-ray, detects theaddress of a detector 21 that detects the γ-ray, and outputs to the dataprocessing device 12 information containing a data of the detectionenergy of the detected γ-ray (information on the detection energyvalue), a data of the detection time (detection time information), and adetector ID (detector address information) corresponding to the addressof the above-described detector 21, as a packet data (information ondetected radiation).

As shown in FIG. 2, the data processing device 12 includes anon-illustrated storage device, a simultaneous measurement device 12A,and a tomogram information preparation device 12B. The data processingdevice 12 captures a packet data (information on detected radiation).The simultaneous measurement device 12A carries out simultaneousmeasurement based on the packet data, in particular, a data of thedetection time, and the detector ID. Then, the simultaneous measurementdevice 12A identifies a detection position of a γ-ray of 511 keV andstores the same into the storage device. The tomogram informationpreparation device 12B prepares functional images based on theidentified position, and displays the same on the display device 13 a.

Incidentally, the test object P is given fluoro-deoxy-glucose (FDG)containing radiopharmaceutical, e.g., ¹⁸F with a half-life period of 110minutes. From the body of the test object P, a pair of 511 keV γ-rays(annihilation γ-rays) are emitted in approximately 180° direction withrespect to each other at the time of annihilation of a positron emittedfrom FDG.

At this time, each detector 21 of the camera 11 surrounds thecircumference of the bed 14. From the detector unit 2 to the dataprocessing device 12, the information on the detection energy value, thedetection time information, and the detector ID obtained based on aγ-ray detection signal (radiation detection signal) generated when thedetector 21 interacts with a γ-ray are outputted for each detector 21contained in the detector unit 2.

As shown in FIG. 3A, 60 to 70 pieces of detector units 2 are removablydisposed in opening portions 11 b of the camera 11 in thecircumferential direction so that the inspection and maintenance may befacilitated. The detector unit 2 is mounted via a unit supporting member2A. Moreover, as shown in FIG. 3B, the detector unit 2 is mounted to thecamera 11, with one end thereof being supported on the unit supportingmember 2A. The unit supporting member 2A is hollow disc-shaped(doughnut-shaped) and comprises a large number (corresponding to thenumber of detection units 2 to be mounted) of opening portions 11 b, toeach of which the detector unit 2 is mounted, in the circumferentialdirection of the camera 11. In order to support the detector unit 2 atone end in this manner, a flange part serving as a stopper is providedat a near side in the body axis direction of an enclosure 30 of thedetector unit 2.

Incidentally, when attempting to arrange the detector units 2 as denselyas possible in the circumferential direction, a circumferentially innerflange portion interferes. Then, this interfering flange portion may beeliminated from the enclosure 30 to leave a circumferentially outerflange portion. Another unit supporting member 2A may be set in the bodyaxis direction so that the both ends of the detector unit 2 may be heldwith the both unit supporting members 2A.

In addition, when mounting the detector unit 2 to the camera 11, a cover11 a is removed to expose the unit supporting member 2A, so that thedetector unit 2 may be inserted and mounted from the opening portion 11b until the flange part bumps against the unit supporting member. Byinserting and mounting in this manner, a connection between therespective connectors of a non-illustrated power supply and signalwiring of the camera 11 and the corresponding connector of the detectorunit 2 is carried out, so that the connection of signals and powersupply between the camera 11 and the detector unit 2 is made.

The configurations of the detector 21, the coupling substrate 20, andthe detector unit 2 will be described in detail later.

Hereinafter, the features of the present embodiment is described.

(Semiconductor Radiation Detector)

First, the detector 21 applied to the present embodiment is describedwith reference to FIG. 4 and FIG. 5. FIG. 4 is a view conceptuallyshowing a laminated structure of the detector 21, the (a) of FIG. 5 is aperspective view schematically showing detection elements andelectrodes, which are component parts of the detector, and the (b) ofFIG. 5 is a perspective view after laminating and integrating these.

As shown in FIG. 4, the detector 21 has a laminated structure whereinboth sides of a semiconductor radiation detection element (hereinafter,referred to as a detection element) 211 composed of a plate-likesemiconductor material S are covered with thin plate-like (filmy)electrodes (anode A, cathode C) and this resulting semiconductorradiation detection element is laminated in five layers, for example.

Among these, the semiconductor material S is composed of a singlecrystal of any one of the above-described CdTe (cadmium telluride),CdZnTe (cadmium zinc telluride), HgI₂ (mercury iodide), TlBr (thalliumbromide), GaAs (gallium arsenide), and the like.

Moreover, for the electrode A (anode A) and the electrode C (cathode C),a material of any one of Pt (platinum), Au (gold), In (indium), and thelike is used.

In addition, in the following descriptions, assume that thesemiconductor material S is a sliced single crystal of CdTe. Moreover,assume that radiation to be detected is a γ-ray of 511 keV used in thePET apparatus 1.

The thickness of one layer of the semiconductor material S (detectionelement 211) shown in FIG. 4 ranges from 0.5 to 1.5 mm, for example. Thethicknesses of the anode A and cathode C are approximately 20 μm,respectively.

In the detector 21 of a laminated structure shown in FIG. 4, the anodesA are connected to each other in common and the cathodes C are connectedto each other in common, so each layer will not detect radiationindependently of other layers. In other words, when a γ-ray interactswith the semiconductor material S, the detector 21 will not discriminatethe occurrence of the interaction in the uppermost layer from theoccurrence of an interaction in the lowermost layer, and the like. Ofcourse, the detector 21 may be configured so as to detect an interactionfor each layer.

Incidentally, the reason why such five layer structure is adopted is asfollows. It is convenient as a detector if the thickness of thesemiconductor material S made thin, because both the rising speed andpulseheight value of a γ-ray detection signal are increased. However, ifthe thickness of the semiconductor material S is thin, there will bemore γ-rays passing straight through the semiconductor material Swithout interacting therewith. For this reason, it is intended to, whileincreasing the charge collection efficiency, reduce the quantity ofγ-rays passing straight through the semiconductor material S and thusincrease the interaction between the semiconductor material S andγ-rays, i.e., aiming at increasing the number of counts.

The configuration of the detectors 21 with such a laminated structurecan provide a more excellent rising characteristic of a γ-ray detectionsignal and a more accurate pulseheight value, and at the same time canincrease the number of counts of γ-rays interacting with thesemiconductor material S, i.e., can increase the sensitivity, as well.

The area of the electrode (anode A, cathode C) preferably ranges from 4to 120 mm². An increase of the area increases the capacitance (straycapacitance) of the detector 21, and this increase of the straycapacitance increases the noise, so it is better if the area of theelectrode is small as much as possible. Moreover, charges generated atthe time of detection of a γ-ray are partly stored in the straycapacitance, so if the stray capacitance increases, problems will occurthat the charge amount stored in a preamplifier 24 a of an analog ASIC24 (see FIG. 7), and in turn the output voltage (pulseheight value)decreases. When CdTe is used as the detector 21, the relative dielectricconstant thereof is 11, and if the area of the detector 21 is set to 120mm² and the thickness thereof is to 1 mm, the capacitance becomes 12 pF,which can not be ignored considering that the stray capacitance of theconnector of the circuit and the like is several pF. Accordingly, thearea of the electrode is preferably set equal to or less than 120 mm².

Moreover, the lower limit of the area of the electrode is determinedfrom the positional resolution of the PET apparatus 1.

In addition, although the semiconductor material S interacting with aγ-ray is assumed to be CdTe in the above description, not to mentionthat the semiconductor material S may be CdZnTe, TlBr, GaAs, or thelike.

(Coupling Substrate; Detector Substrate and ASIC Substrate)

Next, the detailed structure of the coupling substrate (unit substrate)20 mounted in the detector unit 2 is described using FIG. 6. The (a) ofFIG. 6 is a front view showing the coupling substrate, and the (b) ofFIG. 6 is the side view of the (a) of FIG. 6.

The coupling substrate 20 is constructed by connecting a detectorsubstrate 20A, on both sides of which a plurality of detectors 21 aremounted, to an ASIC substrate 20B, on both sides of which a capacitor22, a resistor 23, an analog ASIC 24, and an analog-to-digital converter(hereinafter, referred to as ADC) 25 are mounted and on one side ofwhich a digital ASIC 26 is mounted, via a connector C1.

(Detector Substrate)

As shown in the (a) of FIG. 6, in the detector substrate 20A, on oneside of a substrate body 20 a, for example, 16 detectors 21 are disposedin a horizontal row in the (a) of FIG. 6 corresponding to the body axisdirection of the test object P, and furthermore, four rows of detectors21 are disposed in the vertical direction in the (a) of FIG. 6corresponding to the radial direction with respect to the body axis ofthe test object P, namely, a total of 64 detectors 21 (16 horizontally×4vertically) are disposed in a grid pattern. Moreover, as shown in the(b) of FIG. 6, the detectors 21 are disposed similarly on the othersurface of the detector substrate 20A, as well, and thus a total of 128detectors 21 on both surfaces are disposed in one detector substrate20A.

Here, the more the number of detectors 21, the more easily a γ-ray willbe detected and also the positional accuracy at the time of detection ofa γ-ray can be increased. Accordingly, the detectors 21 are disposed asdensely as possible on the detector substrate 20A.

Incidentally, in the (a) of FIG. 6, in the case where a γ-ray emittedfrom the test object P on the bed 14 travels from the lower side to theupper side of the view (in the direction of an arrow 32, i.e., in theradial direction of the camera 11), it is more preferable that thearrangement of the detectors 21 in the horizontal direction on thedetector substrate 20A be made dense, because the number of γ-rayspassing straight (the number of γ-rays passing through a gap between thedetectors 21) can be reduced. Accordingly, the detection efficiency ofγ-rays can be increased and thus the spatial resolution in the body axisdirection of images obtained can be increased.

In addition, because the detectors 21 are disposed on both sides of thesubstrate body 20 a as shown in the (b) of FIG. 6, the detectorsubstrate 20A of the present embodiment can be commoditized by mountingthe substrate body 20 a on both sides thereof, rather than in the casewhere the detectors 21 are disposed on only one side. Accordingly, thenumber of substrate bodies 20 a can be reduced by half, and thedetectors 21 can be arranged more densely in the circumferentialdirection of the body axis of the test object P. At the same time, asdescribed above, the number of detector substrates 20A (couplingsubstrate 20) can be reduced by half.

In the above description, in the camera 11, horizontally 16 detectors 21are disposed in the body axis direction of the test object P, but notlimited thereto. For example, in the camera 11, horizontally 16detectors 21 may be disposed in the circumferential direction withrespect to the body axis of the test object P.

Moreover, in the detector 21, the surfaces of the electrodes A and Cshown in the (b) of FIG. 5 may be arranged in parallel to the surface ofthe substrate body 20 a, or the surfaces of the electrodes A and C maybe arranged perpendicular to the surface of the substrate body 20 a.

For the purpose of collecting charges, a potential difference (voltage)of 500 V is applied between the anode A and cathode C of each detector21 by means of a high voltage power supply 27 (see FIG. 8), for example.This voltage is supplied from the ASIC substrate 20B side to thedetector substrate 20A side via the connector C1 (see the (a) of FIG.6). Moreover, a γ-ray detection signal outputted when each detector 21detects a γ-ray is supplied to the ASIC substrate 20B side via theconnector C1. For this reason, in the substrate body 20 a of thedetector substrate 20A, there are provided non-illustrated on-boardwiring (used for voltages applied to the detector, used for signaltransfer) that connects the connector C1 to each detector 21. Thison-board wiring has a multilayer structure. The detector substrate 20Aincludes the connector C1 connected to the on-board wiring to beconnected to each detector 21, and is connected to a connector C1 of thelater-described ASIC substrate 20B.

(ASIC Substrate)

Next, the ASIC substrate 20B having ASICs mounted thereon is describedwith reference to FIG. 6. As shown in the (a) of FIG. 6, in the ASICsubstrate 20B, two analog ASICs 24 are mounted on both sides of asubstrate body 20 b, respectively, and one digital ASIC 26 is mounted onone side. In other words, one ASIC substrate 20B includes a total offour analog ASICs 24 and one digital ASIC 26.

Moreover, the ASIC substrate 20B includes 16 ADCs 25 on one side of thesubstrate body 20 b, respectively, i.e., a total of 32 ADCs 25.Moreover, the capacitors 22 and resistors 23 of a number correspondingto the number of detectors 21 are mounted on both sides of one substratebody 20 b. Moreover, in order to electrically connect these capacitors22, resistors 23, analog ASICs 24, ADCs 25, and digital ASIC 26 to eachother, non-illustrated on-board wiring is provided on the ASIC substrate20B (substrate body 20 b) as in the above-described detector substrate20A. This on-board wiring has also a laminated structure.

The arrangement and on-board wiring of each of these circuit elements22, 23, 24, 25, and 26 are made in such a manner that signals suppliedfrom the detector substrate 20A may be supplied to the capacitor 22,resistor 23, analog ASIC 24, ADC 25, and digital ASIC 26 in this order.

In addition, the ASIC substrate 20B includes: the connector C1 formaking electrical connection with the detector substrate 20A, theconnector C1 being connected to the on-board wiring connected to eachcapacitor 22; and a substrate connector C2 for making electricalconnection with the data processing device 12 side (the later-describedunit integrating FPGA side).

(Connection Configuration between Detector Substrate and ASIC Substrate)

As shown in the (b) of FIG. 6, in the detector substrate 20A and theASIC substrate 20B, overlap portions that mutually overlap are providednear the end portions to connect the mutual connectors C1 existing inthese overlap portions. This connection is removably (separably andconnectably) made with fastening screws or the like.

In addition, the reason why such connection is made is as follows. Ifthe coupling substrate 20, to which the detector substrate 20A and theASIC substrate 20B are connected (coupled), is horizontally supported atone end (cantilevered suspension) or supported at both ends, then aforce deflecting or bending this coupling substrate 20 downward will acton the center portion (connection portion) of the coupling substrate 20.Accordingly, if the connection portion is made by abutting the end facesto each other, the connection portion will easily deflect or easilybend. For this reason, it is intended to increase the strength of theconnection portion.

As the connector C1, for example, a spiral contact (registeredtrademark) is used in order to provide excellent electrical connection.The spiral contact (registered trademark) has a characteristic that aball-like connection terminal is in contact with the spiral contact on awide area, thus achieving an excellent electrical connection.

In addition, if the ball-like connection terminal is provided in theASIC substrate 20B side, the spiral contact is provided on the detectorsubstrate 20A side, while if the ball-like connection terminal isprovided in the detector substrate 20A side, the spiral contact isprovided on the ASIC substrate 20B side.

By using such electrical connection structure between the detectorsubstrate 20A and the ASIC substrate 20B, signals can be transmittedfrom the detector substrate 20A to the ASIC substrate 20B at a low loss.If the loss decreases, the energy resolution as the detector 21 will beimproved, for example.

Although in the above-described structure one detector substrate 20A isconnected to the ASIC substrate 20B, the detector substrate may bedivided into multiple portions. For example, horizontally eight andvertically four detectors 21 may be mounted on one substrate to connecttwo detector substrates to the ASIC substrate.

Moreover, the detector substrate 20A and the ASIC substrate 20B may beconstructed with one through-substrate body.

Next, the configuration and function of each circuit element on the ASICsubstrate will be described with reference to FIG. 7 to FIG. 9.

FIG. 7 is a block diagram of the detector unit showing a schematicconfiguration of each of the analog ASIC and digital ASIC, and aconnection relation between the analog ASIC and the digital ASIC.

(Analog ASIC)

FIG. 8 is a block diagram schematically showing the functionalconfiguration of the analog ASIC.

As shown in FIG. 8, one analog ASIC 24 comprises, for example, 32 setsof analog signal processing circuits 33 each includes: a chargesensitive preamplifier (hereinafter, referred to as a preamplifier) 24a; and a fast system 24A and two slow systems (a first slow system 24B,a second slow system 24C) connected to the charge sensitivepreamplifier. One analog ASIC 24 is an LSI integrating 32 sets of analogsignal processing circuits 33.

The analog signal processing circuit 33 is provided for each detector21, and one analog signal processing circuit 33 is connected to onedetector 21. Here, the fast system 24A comprises: a comparator 24 b thatoutputs a timing signal VT for identifying a detection time of a γ-raybased on a γ-ray detection signal VB outputted from the preamplifier 24a; and a threshold control circuit 24 c that sets a threshold voltagevalue VLD inputted to the comparator 24 b. Moreover, in the first slowsystem 24B, for the purpose of calculating a detection energy(pulseheight value) VE1 of a γ-ray based on the γ-ray detection signalVB, a waveform shaper circuit 24 d with a predetermined time constantand a peak hold circuit 24 e are provided and connected in this order.Similarly, in the second slow system 24C, for the purpose of calculatinga pulseheight value VE2 of the amount of electron contribution of thedetection energy (pulseheight value) VE1 of a γ-ray based on the γ-raydetection signal VB, a waveform shaper circuit 24 f with a predeterminedtime constant shorter than that of the waveform shaper circuit 24 d anda peak hold circuit 24 g are provided and connected in this order.

Incidentally, since it takes a certain amount of processing time for theslow systems 24B, 24C to calculate the pulseheight values, these arenamed as “slow”. A signal outputted from the detector 21 and passingthrough the capacitor 22 is amplified by the preamplifier 24 a, and isoutputted as the γ-ray detection signal VB, and is further amplified bythe waveform shaper circuits 24 d, 24 f, and is inputted to the peakhold circuits 24 e, 24 g as the signals VS1, VS2, respectively. The peakhold circuit 24 e holds a maximum value of the waveform-shaped γ-raydetection signal VS1 as described later, i.e., the pulseheight value VE1proportional to an energy value of the detected γ-ray. The peak holdcircuit 24 g holds the pulseheight value VE2 of the waveform-shapedγ-ray detection signal VS2 corresponding to the amount of electroncontribution of the γ-ray detection signal VB as described later.

In addition, although the capacitor 22 and resistor 23 may be providedinside the analog ASIC 24, in the present embodiment, in order to obtainan appropriate capacitor value and appropriate resistance value, and forreasons of reducing the size of the analog ASIC 24, the capacitor 22 andresistor 23 are mounted outside the analog ASIC 24.

Incidentally, the variation in the individual capacitor value orresistance value is assumed small if the capacitor 22 and resistor 23are provided outside the analog ASIC 24.

The output of the first slow system 24B of the analog ASIC 24 issupplied to ADC 25A, and the output of the second slow system 24C issupplied to ADC 25B (see FIG. 8). Furthermore, the output of the fastsystem 24A of the analog ASIC 24 and the outputs of ADCs 25A, 25B aresupplied to the digital ASIC 26.

Here, each of the analog ASICs 24 and the digital ASIC 26 is connectedby means of 32 pieces of wirings each transmitting each of the signalsof 32 channels of fast systems 24A (see FIG. 7). Moreover, the analogASIC 24 and each of the ADCs 25A, 25B are connected to each other, andeach of the ADCs 25A, 25B and the digital ASIC 26 are connected to eachother, respectively, with one piece of wiring that puts together andtransmits the signals of the slow systems 24B, 24C of eight channels ofdetectors 21, (see FIG. 7).

Furthermore, an ADC control signal line for simultaneously controlling apair of ADCs 25A, 25B that are associated corresponding to eightchannels of analog signal processing circuits 33; and a peak holdcontrol signal line obtained by putting eight channels of peak holdcontrol signal lines together into one line, the eight channels of peakhold control signal lines simultaneously controlling the peak holdcircuits 24 e, 24 g of one analog signal processing circuit 33 of eachanalog ASIC 24, are routed from the digital ASIC 26 and connected to theADCs 25A, 25B and the analog signal processing circuit 33.

In addition, each of the above-described eight channels of peak holdcontrol signal lines may be routed from the digital ASIC 26 to eachanalog signal processing circuit 33 without putting them together intoone line.

(Digital ASIC)

Next, the digital ASIC 26 is described with reference to FIG. 7 and FIG.9.

FIG. 9 is a block diagram schematically showing the functionalconfiguration of the digital ASIC. As shown in FIG. 7, the digital ASIC26 comprises 16 sets of detection signal processing parts 34, eachincluding eight timing detection parts 35 and one detector control part36, and one data transfer part 37. The digital ASIC 26 is an LSIintegrating these.

All the digital ASICs 26 provided in the PET apparatus 1 receive a clocksignal from a non-illustrated 500 MHz clock generator circuit (crystaloscillator), for example, and operate in a synchronous manner. The clocksignal inputted to each of the digital ASICs 26 is inputted to therespective timing detection parts 35 in all the detection signalprocessing parts 34.

The timing detection part 35 is provided for each detector 21, and thetiming signal VT is inputted from the comparator 24 b of thecorresponding analog signal processing circuit 33 is inputted the timingdetection part 35. The timing detection part 35 determines a detectiontime of a γ-ray based on a clock signal when the timing signal VT isinputted. Because the timing signal VT is based on the signal of thefast system 24A of the analog ASIC 24, a time close to the truedetection time can be made a detection time (detection timeinformation).

The detector control part 36 comprises an address calculation part 36 a,a detection time correcting part 36 b, a detection energy correctingpart 36 c, a packet data generation part 36 d, a noise determinationpart 36 e, a noise counting part 36 f, and a detector output signalprocessing control part 36 g (hereinafter, referred to as a controlpart).

Upon receipt of detection time information on a detected γ-raycorresponding to the timing signal VT from the timing detection part 35,the address calculation part 36 a identifies the relevant detector IDand outputs this detector ID to the detection time correcting part 36 b,the detection energy correcting part 36 c, the packet data generationpart 36 d, and the noise determination part 36 e. In other words, theaddress calculation part 36 a stores the detector ID for each timingdetection part 35 connected to the address calculation part 36 a inadvance, so that when detection time information is inputted from acertain timing detection part 35, the address calculation part 36 a canidentify a detector ID corresponding to this timing detection part 35.This is possible because the timing detection part 35 is provided foreach detector 21.

Furthermore, after receiving the time information, the addresscalculation part 36 a outputs a peak hold control signal to the analogsignal processing circuit 33 including the above-described identifieddetector ID, and also outputs the detector ID and the ADC control signalto ADCs 25A, 25B.

In addition, the address calculation part 36 a includes anon-illustrated nonvolatile memory and stores therein a detector IDdetermined as abnormal, which is described late. If a detector IDidentified by the timing signal VT coincides with the detector IDdetermined as abnormal, the address calculation part 36 a will notoutput the peak hold control signal and the ADC control signal.

The peak hold circuits 24 e, 24 g of the analog signal processingcircuit 33 that received the peak hold control signal will carry outpeak hold processing to the signals inputted from the waveform shapercircuits 24 d, 24 f. Then, upon receipt of a reset signal from theaddress calculation part 36 a after a predetermined time elapsed, thepeak hold circuits 24 e, 24 g cancels the peak hold processing. ADCs25A, 25B convert the pulseheight values (voltage values) VE1, VE2outputted from the peak hold circuits 24 e, 24 g of the analog signalprocessing circuit 33 corresponding to the detector ID inputted from theaddress calculation part 36 a, into a digital signal and output the sameto the noise determination part 36 e.

The noise determination part 36 e includes a non-illustrated nonvolatilememory and stores therein a correlation data between two pulseheightvalues VE1, VE2 used for determining whether a relevant detection signalis a noise or not. Thus, the noise determination part 36 e determineswhether the relevant signal is a noise signal or a γ-ray detectionsignal based on the inputted pulseheight values VE1, VE2 (the detailswill be described later). If determined as a γ-ray detection signal, thenoise determination part 36 e will not output a noise count signal butoutput the pulseheight values VE1, VE2 to the detection energycorrecting part 36 c.

The detection energy correcting part 36 c includes a non-illustratednonvolatile memory and stores therein each correction value of the gainand offset of each detector 21 and analog ASIC 24 based on thecalibration data collected in advance. Then, using the above-describedcorrection value corresponding to the detector ID inputted from theaddress calculation part 36 a, the detection energy correcting part 36 ccalculates pulseheight value VE1′ and VE2′ that are corrected from thepulseheight values VE1, VE2 and outputs the same to the detection timecorrecting part 36 b. Moreover, the detection energy correcting part 36c generates, based on the corrected pulseheight value VE1′ and VE2′, theinformation on the detection energy value corresponding to an energy ofthe detected γ-ray, and outputs the same to the packet data generationpart 36 d.

If determined as a noise signal, the noise determination part 36 eoutputs a noise count signal to the noise counting part 36 f along withthe detector ID, and outputs a reset signal to the packet datageneration part 36 d but does not output the pulseheight values VE1, VE2to the detection energy correcting part 36 c.

The noise counting part 36 f includes a non-illustrated nonvolatilememory or volatile memory and counts and stores therein the noise countfor each detector ID. If the noise count is equal to or greater than aspecified reference value, the noise counting part 36 f determines arelevant detector 21 as abnormal and outputs this abnormalitydetermination to the control part 36 g. Moreover, the control part 36 gincludes a non-illustrated nonvolatile memory and stores therein theabnormality determination output from the noise counting part 36 g.Based on this, the control part 36 g controls data processing in thedetector control part 36 with respect to an output signal from thedetector 21 corresponding to the detector ID that is determined asabnormal.

The detection time correcting part 36 b corrects the detection timeinformation inputted from the address calculation part 36 a based on thepulseheight values VE1′, VE2′ inputted and corrected from the detectionenergy correcting part 36 c, and outputs the same to the packet datageneration part 36 d.

The packet data generation part 36 d appends the corrected detectiontime information and the detector ID to the information on the detectionenergy value from the detection energy correcting part 36 c, and therebygenerates a packet data (information on the detected γ-ray, informationon the detected radiation), which is digital information, and outputsthe same to the data transfer part 37. The data transfer part 37, forexample, periodically transmits the packet data outputted from thepacket data generation part 36 d of each detection signal processingpart 34 to the unit integration FPGA 31 (Field Programmable Gate Array,hereinafter, referred to as FPGA) provided outside the enclosure 30 ofthe detector unit 2 (see FIG. 15, FIG. 16) that houses twelve couplingsubstrates 20 therein. FPGA 31 transmits these digital information tothe data processing device 12 via information transmission wiringconnected to a connector 38.

In addition, the method for correcting detection time information in thedetection time correcting part 36 b, the method for determining whethera relevant detection signal is a γ-ray detection signal or a noise inthe noise determination part 36 e, and the detailed operations of thenoise counting part 36 f and control part 36 g will be described later.

(Detector Unit; Unitization by Housing Coupling Substrates)

Next, unitization by housing the above-described coupling substrates 20into the enclosure 30 is described.

As shown in FIG. 10, the detector unit 2 comprises twelve couplingsubstrates 20, a high voltage power supply device PS that supplies acharge collecting voltage to these twelve coupling substrates 20, FPGA31, a signal connector that transfers signals to/from the outside, theenclosure 30 (see FIG. 11) that houses or holds a power supply connectorand the like for receiving power supply from the outside, and the like.

As shown in FIG. 10 and FIG. 11, in the enclosure 30, three rows ofcoupling substrates 20 are housed so as not to overlap to each other inthe depth direction (in the body axis direction of the test object P),while four coupling substrates 20 are housed side by side in thecircumferential direction of the camera 11. Namely, twelve couplingsubstrates 20 are housed in one enclosure 30. In order to house thisway, a guide member 39 including four rows of guide grooves G1 suitablyspaced apart to each other in the circumferential direction, the guidegroove G1 extending in the depth direction, is attached to a top endportion of the enclosure 30. The guide member 39 includes an opening 40at a position opposite to each connector C3 of a top plate 30 a in eachguide groove G1 portion.

Furthermore, four guide members 41 each including a guide groove G2extending in the depth direction are attached to the upper surface of abottom plate 30 b of the enclosure 30, the guide members 41 beingsuitably spaced apart to each other in the circumferential direction(see FIG. 11). The guide grooves G1, G2 have a depth enough to housethree coupling substrates 20. An edge portion on the ASIC substrate 20Bside of the coupling substrate 20 is inserted into the guide groove G1,and an edge portion on the detector substrate 20A side of the couplingsubstrate 20 is inserted into the guide groove G2. Three couplingsubstrates 20 are to be held side by side in the depth direction of theguide grooves G1, G2.

Incidentally, in the coupling substrate 20, an edge portion on the ASICsubstrate 20B side and an edge portion on the detector substrate 20Aside are adapted to slide in the guide grooves G1, G2, respectively, sothat the coupling substrate 20 can be easily slid in the guide groovesG1, G2 using fingers or the like to position at a predetermined place.At this time, each of the substrate connectors C2 is positioned in theopening 40 portion. After a predetermined number of coupling substrates20 are disposed in the enclosure 30, the top plate 30 a is removablymounted to the top edge of the enclosure 30 with screws or the like.Each connector C3 provided in the top plate 30 a is inserted into thecorresponding opening 40 and connected to the corresponding substrateconnector C2.

In addition, the upper and the lower of the enclosure 30 are referred tothe positional relation when the enclosure 30 is removed from the camera11, and as shown in FIG. 2, when the enclosure 30 is provided in thecamera 11, the upper and the lower will be reversed, or the upper andthe lower will be rotated by 90° to be at a horizontal position or at adiagonal position.

As shown in FIG. 11, the top plate 30 a of the enclosure 30 is providedwith the FPGA 31 and the connector 38 (see FIG. 10) in addition to theabove-described four rows of guide grooves G1. The connector 38 isconnected to the FPGA 31. Programming can be done using FPGA 31 on site.FPGA 31 differs in this point from ASIC with which programming cannot bedone. Accordingly, as in the present embodiment, FPGA 31 canappropriately respond to, for example, even the case where the numberand type of the coupling substrates 20 to be housed have changed, oralso the case where the number of coupling substrates has changed.

In addition, since the detector 21 used in the present embodiment,wherein CdTe is used as the semiconductor material S, will generatecharges in response to light, the enclosure 30 is formed of a materialhaving light shielding characteristics, such as aluminum or an aluminiumalloy, and is thus adapted to eliminate a gap through which lightenters. That is, the enclosure 30 is constructed so as to have lightshielding characteristics. Incidentally, if the light shieldingcharacteristics are secured by other means, the enclosure 30 itself doesnot need to have light shielding characteristics and just needs to be aframe that removably holds the detector 21. For example, the enclosure30 may be of a frame structure, thus eliminating a need for alight-shielding face plate, and the like.

Moreover, packet data (all the packet data for all the detectors 21 ofall the coupling substrates 20) outputted from the data transfer parts37 of all the coupling substrates 20 in the detecting unit 2 is sentfrom FPGA 31 provided in the detecting unit 2 to the data processingdevice 12.

(Power Supply)

Next, the high voltage power supply device PS for supplying a chargecollecting voltage is described. As shown in FIG. 10, in the detectorunit 2, the high voltage power supply device PS for supplying a chargecollecting voltage to each detector 21 is mounted in a space formed by apartition 30 c composed of a conductor metal material inside theenclosure 30 on the rear surface side of FPGA 31. This high voltagepower supply device PS is supplied with a low voltage power, and isadapted to boost up this voltage to 500 V using a non-illustratedvoltage boosting type DC-DC converter, and to supply the same to eachdetector 21. Incidentally, 64 detectors 21 are provided on one side and128 detectors 21 are provided on both sides per one detector substrate20A. Then, twelve coupling substrates 20 are housed in one enclosure 30.Accordingly, the high voltage power supply device PS supplies a voltageto 128×12=1536 detectors 21.

In the present embodiment, the high voltage power supply device PScontained in the detector unit 2 is connected to an external low voltage(5 to 15 V) DC power supply by power wiring via a power connector 42 andconnector 38 provided in the top plate 30 a. High voltage side terminalpins of the high voltage power supply device PS are connected to twelveconnectors C3 provided on the top plate 30 a, respectively, via theconnector 43 provided on the top plate 30 a, by means of high voltagepower supply wiring 44. In FIG. 10, only one high voltage power supplywiring 44 is shown as to be connected from the connector 43 to theconnector C3 by way of illustration, but the high voltage power supplywiring 44 is actually routed from the connector 43 to each connector C3.

The high voltage power supply is connected to the electrode C of eachdetector 21 provided in the substrate body 20 a, respectively, via theconnector C3, the connector C2 of each coupling substrate 20, thenon-illustrated high voltage power supply connector C1 in the substratebody 20 b, and non-illustrated high voltage power wiring in thesubstrate body 20 a. The connectors C1, C2 include a connector used forhigh voltage power wiring in addition to the connectors for transmittingthe output signals of the detector 21.

Incidentally, a voltage supplied from the connector 38 to the highvoltage power supply device PS is boosted up to 500 V by anon-illustrated DC-DC converter inside the high voltage power supplydevice PS, and after the boost up, this is supplied, through theinterior of the top plate 30 a of the enclosure 30, to the ASICsubstrate 20B to the detector substrate 20A to each detector 21 for eachcoupling substrate 20. That is, the enclosure 30 (top plate 30 a)includes the voltage supply wiring for supplying a voltage from the highvoltage power supply device PS to each coupling substrate 20. Moreover,each coupling substrate 20 includes the voltage supply wiring forsupplying each detector 21 with a voltage supplied from the high voltagepower supply device PS via the substrate connector C2.

In addition, the high voltage power supply device PS may be directlyconnected to the high voltage power wiring provided in the substratebody 20 a via a connector, instead of via the top plate 30 a. Moreover,the high voltage power connector may be disposed separately from theoutput signal connector of the detector 21.

<<Operational Description of Analog ASIC and Digital ASIC>>

Next, processing to the detector output signal, the processing being afeature of the present invention, will be described with reference toFIG. 12 to FIG. 16.

(Operational Description of Analog ASIC)

FIG. 12A is a view, where the detector 21 is modeled with a singlelayer, showing a portion where the comparator 24 b outputs the timingsignal VT, which is a voltage signal, based on an output current pulseIA from the detector 21.

FIG. 12B is a view showing a relation among the waveform of the currentpulse IA generated at the input side of the preamplifier 24 a in FIG.12A, the voltage signal VB obtained as an output by converting thiscurrent pulse IA into a voltage signal in the preamplifier 24 a, i.e.,the waveform of the voltage signal VB appearing at the input side of thecomparator 24 b, and the timing signal VT that is outputted from thecomparator 24 b based on this voltage signal VB with a predeterminedvoltage VLD being as a threshold.

In the detector 21, the both electrodes of cathode C and anode A areconnected to the high voltage power supply 27 and to a resistor 23 via acapacitor 22, respectively, and the anode A on the resistor 23 side isconnected to the input side of the preamplifier 24 a. The output side ofthe preamplifier 24 a is connected to an input side (+) of thecomparator 24 b, and the output voltage VLD from a threshold controlcircuit 24 c is connected to a reference side (−) of the comparator 24b, and if the output voltage VB becomes equal to or greater than thepredetermined voltage threshold VLD, the timing signal VT appearing atthe output terminal of the comparator 24 is inverted.

If a γ-ray is incident upon the detector 21 and absorbed, anelectron-hole pair corresponding to the absorbed γ-ray energy will begenerated, and the electron and hole are induced by an electric fieldapplied by the high voltage power supply 27, and thereby the electronwill transfer to the anode A side and the hole will transfer to thecathode C side. This transfer of the electron and the hole results inthe current pulse IA occurring at the input side of the preamplifier 24a. Since the detector 21 has a finite size, the output signal waveformof the current pulse IA varies depending on a position where a γ-ray isabsorbed. This is because if a γ-ray is incident upon the vicinity ofthe cathode C of the detector 21, the ratio for the generated electronsto contribute to the current pulse IA (hereinafter, referred to as a“contribution of electron”) is large and on the contrary if thegenerated electrons are absorbed in the vicinity of the anode A, theratio for the generated electrons to contribute to the current pulse IAis small. Since the mobility (transfer speed) of an electron istypically 10 times or more the mobility of a hole, the waveform (outputsignal waveform) of the current pulse IA outputted by the detector 21will vary depending on the position where the electrons and holes areabsorbed in the detector 21. If the distance between the anode A andcathode C is denoted by L0 and the distance between the anode A and aplace where an electron-hole pair is generated is denoted by Lx, then inthe case of Lx/L0=1, the detector 21 detects an energy approximatelyequal to the current resulting from the transfer of electron. Moreover,in the case of Lx/L0=0.5, the electron and the hole each half contributeto the current. Moreover, in the case of Lx/L0=0, the generated currentwill substantially depends on the transfer of hole.

If the absorption of a γ-ray occurs in the vicinity of the cathode C ofthe detector 21, then the voltage signal VB will be of a signal waveformwith the contribution of electron approximately equal to one (100%),thus resulting in the earliest timing detection. Moreover, if theabsorption of a γ-ray occurs between the anode A and cathode C to causethe contribution of electron equal to 0.5 (50%), the comparator 24 bwill output the voltage signal VB rising at a timing later than theformer. Moreover, if the absorption of a γ-ray occurs in the vicinity ofthe anode A, the voltage signal VB will be of a signal waveform with thecontribution of electron approximately equal to 0 (0%), thus resultingin the latest timing detection. Depending on the configuration of thedetector 21 or on the value of a voltage applied to the detector 21,generally, when the detector 21 is used in the PET device 1, thetransfer time of the electron (te) in the detector 21 ranges from 20 ns(nanosecond) to 50 ns and the transfer time of the hole (th) in thedetector 21 approximately ranges from 200 ns to 500 ns. In this case,depending on the voltage threshold VLD to be compared with, the timedifference between the earliest timing detection and the latest timingdetection approximately ranges from 200 ns to 300 ns, thus causing a bigproblem in the simultaneous measurement using the semiconductorradiation detector 21.

FIG. 13A is a view showing the scattering and absorption of radiation inthe detector 21, and specifically shows an energy absorption spectrumwhen the scattering occurs. The horizontal axis represents the detectionenergy of γ-rays and the vertical axis represents the frequency ofoccurrence. The detector 21 does not necessarily absorb all the energyof the detected γ-ray. The detector 21 is known to stochastically causeabsorption or scattering statistically depending on the quality of thematerial of elements, shape, and the energy of γ-rays. Since thepulseheight value (voltage value) VE1 which the peak hold circuit 24 eoutputs based on a signal outputted by the detector 21 is proportionalto the γ-ray energy absorbed by the detector 21, this pulseheight valueVE1 is assumed here to represent the detected energy of the γ-ray. Ifthe detected energy VE1 exceeds a certain value, the frequency ofoccurrence of absorption will reach a peak at the certain value (here,511 keV). At this time, the rate of occurrence of scattering is minimal.

FIG. 13B is a view showing a state where the scattering and absorptionof a γ-ray occur in the detector 21, FIG. 13C is a view showing arelation among the waveform of the current pulse IA occurring at theinput side of the preamplifier 24 a in FIG. 13B, the voltage signal VBobtained as an output by converting this current pulse IA into a voltagesignal in the preamplifier 24 a, and the timing signal VT outputted fromthe comparator 24 b based on this voltage signal VB.

When a γ-ray enters the detector 21, the detector 21 generates anelectron-hole pair corresponding to an absorbed energy, and the electronand hole are induced by an electric field generated by an appliedvoltage, and thereby the electron will transfer to the anode A side andthe hole will transfer to the cathode C side. This transfer of theelectron and the hole results in the current pulse IA occurring at theinput side of the preamplifier 24 a. The example here shows the casewhere the energy of the incident radiation is 511 keV, wherein all theenergy is absorbed (on the left in the view) and the half the energy isabsorbed and the remaining half energy is scattered (on the right in theview), at the same position between the anode A and the cathode C. Aboutthe current pulse IA of FIG. 13C, in either case, it takes the same timete for the generated electron and hole to transfer to the anode A andthe cathode C, respectively, however, in the latter case the detector 21outputs the current half the former case.

For example, the waveform of the voltage signal VB at the contributionof electron of 100% is described. When the detected energy changes, forexample, when one energy is 511 keV and the other energy is 255 keV, thenumber of electrons occurring in the vicinity of the cathode C insidethe detector 21 changes, and the both cases show no difference in thetime interval (transfer time of the electron) te of the current pulse IAbut shows a difference in the current value (pulse height), so that inthe waveform of the voltage signal VB at the input side of thecomparator 24 b, the energy of 255 keV shows a smaller gradient than theenergy of 511 keV. Accordingly, when compared with a specified voltagethreshold VLD in the comparator 24 b, the energy of 255 keV delays by atime t2 as compared with the energy of 511 keV.

Returning to FIG. 8, the time constant of the waveform shaper circuit 24d is set to 1000 ns, and the time constant of the waveform shapercircuit 24 f to 50 ns, for example. The electron-hole pair occurring inthe detector 21 is converted into the voltage signal VB by thepreamplifier 24 a, and thereafter the comparator 24 outputs the timingsignal VT obtained by discriminating the voltage signal VB with thepredetermined voltage threshold VLD. The voltage signal VB is alsosubjected to noise elimination processing and waveform shaping throughthe waveform shaper circuit 24 d to be converted into a voltage signalVS1 corresponding to an energy E1 of the detected γ-ray, and then thepeak hold circuit 24 e outputs a pulseheight value VE1 of the voltagesignal VS1. The voltage signal VB is converted into a voltage signal VS2corresponding to an energy E2 of the detected γ-ray, the energy E2 beingcontributed by only the transfer of the electron filtered through thewaveform shaping circuit 24 f with a time constant 50 ns correspondingto the mobility of the electron of an electron-hole pair. Then, the peakhold circuit 24 g outputs a pulseheight value VE2 of the voltage signalVS2.

FIG. 14 is a view showing the transient response waveforms of therespective output signals VB, VT, VS1, VE1, VS2, and VE2 in FIG. 8. Whenthe voltage signal VB reaches or exceeds the voltage threshold VLD, thetiming signal VT changes from the low level to the high level, and thepeak hold circuit 24 e and the peak hold circuit 24 g will hold thepulseheight values of the respective outputs VS1 and VS2 of the waveformshaper circuit 24 d and the waveform shaper circuit 24 f. Theabove-described time constant 1000 ns of the waveform shaper circuit 24d is set to a time constant longer than the transfer time of the holeinside the detector 21 to broaden the pass band, thereby allowing forthe passage of a total of the γ-ray detection signals VB of the amountof electron contribution, the amount of hole contribution, and bothsignal components. On the other hand, the above-described filter timeconstant 50 nsec of the waveform shaper circuit 24 f is set to beequivalent to the transfer time of the electron inside the detector 21to narrow the pass band, thereby preventing the passage of the signalcomponents other than the amount of electron contribution of the γ-raydetection signals VB as much as possible. This configuration makes itpossible to acquire information on the detected energy and thecontribution of electron.

(Operation of Noise Determination Part)

As described above, the voltage signal VB that is outputted from thepreamplifier 24 a when the detector 21 detects a γ-ray depends on theposition of the cathode C and the anode A and has a time-varyingdistribution specific to γ-ray detection signals. For this reason, thereis a correlation between the respective pulseheight value VE1 andpulseheight value VE2 of the output signals outputted from the waveformshaper circuit 24 e with a time constant 1000 ns and the waveform shapercircuit 24 g with a time constant 50 ns, wherein the correlationcorresponds to one point in the regions 81 to 83 in FIG. 15 depending onthe detection energy and the position between the anode A and thecathode C where the absorption occurs inside the detector 21.

FIG. 15 is a view showing a correlation between the pulseheight valueVE1 and the pulseheight value VE2, wherein the pulseheight value VE2 ofthe amount of electron contribution is plotted in the horizontal axisand the pulseheight value VE1 including both contributions of electronand hole is plotted in the vertical axis. In FIG. 15, the region 81 is aregion where the amount of electron contribution is dominant, the region83 is a region where the amount of hole contribution is high, and theregion 82 is a region where a part of the contribution of hole is addedto the contribution of electron.

Then, the region 80 is a region of noise signals. The noise signal doesnot necessarily have a δ function-like spike waveform, but may have abroad waveform in time scale.

The noise determination part 36 e includes a non-illustrated nonvolatilememory and stores therein a correlation table data having thepulseheight values VE1, VE2 as shown in FIG. 16 as parameters inadvance. Based on the correlation table data stored in theabove-described nonvolatile memory, the noise determination part 36 ecarries out determination processing to determine a relevant detectionsignal as a noise signal if the inputted pulseheight values VE1, VE2correspond to the region 80.

In addition, in place of the table data indicative of the correlationshown in FIG. 16, an equation expressing a gradient of the boundarybetween the regions of FIG. 15 may be stored in the above-describednonvolatile memory, but the boundary between the region 80 and a regionputting together the regions 81, 82 and 83, is not a straight line, so atable lookup type is better in order to accurately make a noisedetermination.

When the noise determination part 36 e, based on the inputtedpulseheight values (voltage values) VE1, VE2, determines that the γ-raydetection signal VS1 is a γ-ray detection signal resulting from theintended γ-ray detection, i.e., that the γ-ray detection signal VS1 isnot a noise signal, the noise determination part 36 e will not output anoise count signal but output two pulseheight values VE1, VE2 to thedetection time correcting part 36 b and output the pulseheight value VE1to the packet data generation part 36 d. When the noise determinationpart 36 e determines that the γ-ray detection signal VS1 is a noisesignal, the noise determination part 36 e will output a noise countsignal to the noise counting part 36 f and output a reset signal to thepacket data generation part 36 d, but will not output the pulseheightvalues VE1, VE2 to the detection time correcting part 36 b and will notoutput the pulseheight value VE1 to the packet data generation part 36d.

Accordingly, if determined as a noise signal, the detection timecorrecting part 36 b will not carry out the later-described correctionof detection time information and the packet data generation part 36 dwill not output an unnecessary data to the data processing device 12that carries out simultaneous measurement processing, so that the loadof the digital ASIC 26 is reduced and the load of signal processing onthe downstream side is also reduced.

(Noise Counting Part)

Moreover, the noise counting part 36 f has a nonvolatile memory functionto store, for eight channels of detectors 21, the number of times arelevant detection signal is determined as a noise signal. Upon input ofa noise count signal from the noise determination part 36 e, the noisecounting part 36 f will, based on the inputted detector ID, incrementthe noise count number of the relevant detector ID by one count andstores the same. Then, the noise counting part 36 f checks if the noisecount number exceeds a specified reference value for each one increment,and if the noise count number exceeds the specified reference value thatis set in advance, the noise counting part 36 f makes “abnormalitydetermination” (determines as faulty) and outputs the detector ID andthe abnormality determination to the control part 36 g. The noisecounting part 36 f resets to zero the noise count value for the detectorID determined as abnormal, and then newly starts to count.

Here, the noise counting part 36 f may adopt a method of carrying outnoise counting for each specified period of time and automaticallyresetting to zero count after a specified period of time.

(Operation of Detector Output Signal Processing Control Part)

The control part 36 g includes a nonvolatile memory function. Uponreceipt of the abnormality determination, the control part 36 g storesan abnormality determination time based on a system clock and thedetector ID, and inputs and stores this detector ID to the addresscalculation part 36 a.

Even if the timing signal VT of a channel of the detector 21corresponding to the stored detector ID is inputted, the addresscalculation part 36 a will not carry out address calculation processingand will not output the detector ID, thereby preventing the dataprocessing in the detector control part 36. Namely, since the addresscalculation part 36 a will not output an ADC control signal to theabnormal detector 21, the pulseheight value will not be read by ADCs25A, 25B and the calculation processing using VE1, VE2 will not becarried out.

Moreover, the control part 36 g outputs to the packet data generationpart 36 d the faulty information containing the detector ID and theinformation on the time when determined as abnormal, and transmits thesame to the data processing device 12 to store this into a storagedevice of the data processing device 12 as a history in which the faultyinformation is put in chronological order. Thus, an operator canperiodically display a distribution of abnormal detectors 21 andchronological changes and inspect the same.

Moreover, concerning the detector ID that has been determined asabnormal, the control part 36 g, based on the last abnormalitydetermination time information which the control part 36 g itselfstores, determines whether or not new abnormality determinationinformation has not been inputted for a predetermined elapsed timeperiod during the operating time of the PET device 1. When theabnormality determination has not been inputted from the noisedetermination part 36 e for the predetermined elapsed time period, thecontrol part 36 g will erase the relevant detector ID that has beendetermined abnormal, the relevant detector ID being previously stored inthe address calculation part 36 a, from the memory. Moreover, thecontrol part 36 g outputs to the packet data generation part 36 d thedetector ID and the information on the time when the abnormalitydetermination is canceled, and transmits the same to the data processingdevice 12 to stores this into a storage device of the data processingdevice 12 as the faulty information.

The detector 21 may temporarily output a noise signal due to dustadhesion, or the like, and then the dust may come off from the detector21 and the detector 21 may not output the noise signal and return to thenormal state. In this case, the control part 36 g resets the informationon the abnormal detector ID, which was stored in the address calculationpart 36 a so that the detector control part 36 subsequently starts toprocess a γ-ray detection signal of a channel of the relevant detector21.

In addition, in place of receiving abnormality determination from thenoise counting part 36 f, the control part 36 g may check the noisecounting part 36 f at a predetermined cycle, and check whether the noisecount exceeds a specified reference value, which is set in advance, tomake abnormality determination and control the reset of the count of thenoise counting part 36 f.

Moreover, the method for stopping the signal processing to the detector21 is not limited to the above-described method. For example, the highvoltage power supply 27 individually supplied to the detector 21 may beturned off, or the address calculation part 36 a may append anabnormality flag, i.e., identification information indicating to thepacket data generation part 36 d that the relevant detection signal isfaulty, for a packet data of the relevant detector ID, instead ofstopping the address calculation. In this case, the data processing part12 in the subsequent stage detects the abnormality flag contained in thepacket data and excludes the relevant packet data so as not to besubjected to the signal processing.

In addition, although the illustration of signal wirings is omitted inFIG. 9, the data, such as the correlation data used in the noisedetermination part 36 e shown in FIG. 15, the reference value used forabnormality determination in the noise counting part 36 f, can bechecked or exchange using the input operation part 13 b of the operatorconsole 13.

(Correction of Detection Time Information)

Next, a method is described, in which time information obtained when thedetector 21 detects a γ-ray is corrected with two slow systems 24B, 24C.

From the information on the contribution of electron described above, acorrection value for the timing signal VT can be calculated by using arelation at which position of the gap between the cathode C and theanode A in the detector 21 shown in FIG. 12B a γ-ray energy is absorbed.

Since the waveform of VB corresponding to each contribution of electroncan be known in advance, it is possible to calculate a time periodbetween a time when VT becomes the high level, i.e., a time when thewaveform of VB corresponding to the relevant contribution of electroncrosses the horizontal line corresponding to VB=VLD, and a time when anevent of detecting a true γ-ray occurs. If the time information, whichis the output of the timing detection circuit 35, is corrected usingthis time period value, the correct time information can be obtained.Moreover, if N pieces of contribution of electron are arbitrarilyselected and then a correction value corresponding to each of them iscalculated in advance and stored in the detection time correcting part36 b as a table, then the correction value does not need to becalculated each time the voltage signals VE1, VE2 are inputted. This iseffective in accelerating the signal processing.

Next, consider the case where the above-described compton scattering hasoccurred in the detector 21. Even in this case, based on the relation ofchanges in the timing signal VT due to the detection energy shown inFIG. 13C, the timing signal VT can be properly corrected to obtain thetime information.

Since the voltage signal VE1 is proportional to the voltage signal VB,the value of the voltage signal VB can be determined by monitoring thevoltage signal VE1. A difference between the voltage signal VB and avoltage value equivalent to the energy inherent in a γ-ray correspondsto the voltage value equivalent to the energy lost due to scattering.Moreover, since the transfer time te of an electron is uniquelydetermined by the element size of the detector 21 and a voltage appliedto the detector 21, the time when the waveform of each voltage signal VBchanges from a monotonic increase to a constant state is determined, andthe gradient of the monotonically increasing portion between twowaveforms can be found. The time t2 can be found as an interval betweentwo points where these waveforms cross the horizontal line correspondingto VB=VLD. If the timing signal VT is corrected using this t2, thecorrect time information can be acquired. Also in this case, similarly,if N pieces of values of the voltage signal VB are arbitrarily selectedand then a correction value corresponding to each of them is calculatedin advance and stored into the detection time correcting part 36 b as atable, then the correction value does not need to be calculated eachtime the voltage signal VE1, VE2 are inputted. This is effective inaccelerating the PET apparatus 1.

FIG. 16 shows a correction data table having the voltage VE1 and voltageVE2 as parameters. The correction data table in FIG. 16 is atwo-dimensional matrix having two-dimensional data as elements, whereinthe two-dimensional data comprise correction values based on theinformation VE2 of the contribution of electron described above andcorrection values based on the total energy information VE1, i.e., theenergy quantity value that decreased due to scattering. Accordingly, ifthe detection time correcting part 36 b includes this correction datatable, a variation in the contribution of electron depending on theposition of γ-ray energy absorption in the radiation detector 21, avariation in the detected energy due to scattering of radiation in thedetector 21, and a deviation from a true value of the timing signal dueto the both variations can be corrected at once, which is effective inaccelerating the PET apparatus 1.

(Effect of Data Processing Device)

Thus, a packet data containing (1) information on the detection energyvalue, (2) detection time information, and (3) detector ID, the packetdata being outputted from the digital ASIC 26, is transmitted to thedata processing device 12 (see FIG. 2) in the subsequent stage via FPGA31 and the wiring used for information transmission. Based on the packetdata transmitted from the digital ASIC 26, the simultaneous measurementdevice 12A of the data processing device 12 carries out simultaneousmeasurement processing to count a pair of γ-ray as one piece, the pairof γ-ray having undergone the simultaneous measurement processing, andidentifies the positions of two detectors 21, which detected the pair ofγ-ray, from their detector IDs. Here, the simultaneous measurementprocessing is referred to the processing to determine, upon detection oftwo γ-rays having a specified energy within a time window of a set time,whether or not these γ-rays are a pair of γ-ray occurring due toannihilation of one positron, and leave only the γ-ray pair as a dataused for image generation.

When there are three or more γ-ray detection signals detected within theabove-described time window (there are three or more detectors 21 thatdetect the γ-rays), the data processing device 12 determines which is aseries of γ-ray detection signals due to the compton scattering involvedin the incidence of one γ-ray of 511 keV among these three or more γ-raydetection signals by using the information on the detection energyvalue, or the like of these γ-ray detection signals, and identifies adetector ID, upon which the γ-ray at this time entered first, andfurther determines whether the γ-rays are a pair of γ-ray of 511 keV,and subsequently carries out the simultaneous measurement processing.

The identified pair of detectors 21 are simultaneously measured togenerate one count. Moreover, the tomogram information preparationdevice 12B of the data processing device 12 prepares an accumulationposition of radiopharmaceutical, i.e., tomogram information on the testobject P at a malignant tumor position, by using the count valueobtained in the simultaneous measurement and the positional informationon the detectors 21. This tomogram information is displayed on thedisplay device 13 a. The information, such as the above-describeddigital information, the count value and the positional information onthe detectors 21 obtained in the simultaneous measurement, the tomograminformation, is stored in a storage device of the data processing device12.

Effects of First Embodiment

The phenomenon that the detection time delays when the position, wherethe energy of a γ-ray is absorbed, and the contribution of electroninside the detector 21 change as described using FIG. 12A and FIG. 12B,can be uniquely determined if the detected γ-ray energy and thecontribution of electron are determined. The same applies to the casewhere the radiation scatters as described using FIG. 13A to FIG. 13C.

Thus, according to the present embodiment, the timing signal VT can becorrected by transmitting to the detector control part 36 theinformation, such as the detection timing VT, the detected γ-ray energy(pulseheight value) VE1, and the pulseheight value VE2 of the amount ofelectron contribution of the detected γ-ray energy. Distributions oftiming detection times differ as follows. Before the correction, thedistribution corresponding to low contribution ratios of electronmobility spreads widely to the later time side, while after thecorrection, observational data samples gather near the true detectiontime. This increases the probability of assuming true coincidentalevents to be the equal time, and thus effective events in theobservational data will increase. As a result, this makes it possible tosuppress a loss of effective data, which is the greatest disadvantage ofthe PET device 1 using the semiconductor radiation detector 21, and toacquire a high contrast and high resolution image in an image pickuptime equivalent to that of the conventional PET apparatus using ascintillator.

Moreover, the present embodiment can reduce the effect of timingvariations due to noise. The noise is generated from the detector 21 orthe preamplifier 24 a, in particular. Because in the conventionaltechnique disclosed in JP-A-2002-243858, the output of a preamplifier issupplied to a comparator via a current control part, the waveformgradient of a γ-ray detection signal VB is made gentle, thus degradingthe ratio (S/N) of signal intensity to noise intensity. In the presentembodiment, there is no path to output the timing signal VT, i.e., noblock such as the current control part having characteristic similar tothat of a filter causing S/N degradation in the fast system 24A. It istherefore possible to carry out comparator operations in an excellentS/N condition and suppress variations in the timing signal VT due tonoise.

Here, because both the pulseheight value VE1 corresponding to a detectedγ-ray energy and the pulseheight value VE2 of the amount of electroncontribution are signals outputted via the filter, they are subjected toprocess variations. However, because these types of variations aresimply caused by element variations inherent in the circuit elementsconstituting the filter, not caused by noise, calibration can correctthe process variations for each input signal.

Moreover, the present embodiment can shorten the image pickup time inthe PET apparatus 1. Because the conventional technique disclosed in theabove-described JP-A-2002-243858 always forcibly adjusts the slew rateof a high-speed timing signal to that of a low-speed timing signal, timevariations due to noise in the comparator to detect timings or due tocircuit offsets may tend to be significant even when a lot of eventswith a high contribution of electron occur. In contrast, the presentembodiment determines a correction value for the timing signal VT usingthe pulseheight value VE2 for observing the contribution of electron, sotiming can be detected using a faster rising waveform for a γ-raydetection signal with a high contribution of electron. Accordingly, onaverage, the present embodiment decreases variations in timing detectionand thereby provides more effective counts than the above-describedconventional technique, thus allowing the image pickup time in the PETapparatus 1 to be shortened.

Moreover, the present embodiment can improve the diagnostic imagecontrast. Increasing time variations also increases chances of an errorthat leads to assume two signals to occur at the same time though theycorrespond to asynchronously occurring events of radiation incidence.Accordingly, the conventional technique disclosed in JP-A-2002-243858may degrade the diagnostic image contrast due to incorrect detection. Incontrast, the present embodiment enables image pickup to acquire moreeffective counts as described above and thus can improve the tomographicimage contrast.

Moreover, according to the present embodiment, in the noisedetermination part 36 e, the pulseheight value VE1 and the pulseheightvalue VE2 are compared with the correlation table data for each outputsignal of the preamplifier 24 a to thereby determine whether therelevant detection signal is an intended γ-ray detection signal or anoise. If determined as a noise in the noise determination part 36 e, adetection time correction processing to the relevant signal andgenerating a packet data and sending this packet data to the dataprocessing device 12 are not allowed to be carried out. Accordingly, thesignal-processing load of the digital ASIC 26 and FPGA 31 in the PETapparatus 1 with a high counting rate can be reduced.

Especially, in the PET apparatus 1, such that a plurality of γ-raydetection signals are determined as which are caused by the intendedincident γ-ray of 511 keV, and are used for PET image generation, whenan incident γ-ray generates the plurality of γ-ray detection signalsacross a plurality of detectors 21, i.e., also when there is scattering,the threshold for a detected γ-ray energy value in the data processingdevice 12 needs to be set lower to carry out signal processing thattakes into account the scattering. In this case, noise is more likely tobe regarded as a γ-ray detection signal and be a target of the signalprocessing. However, the present embodiment makes it possible toefficiently determine and select a γ-ray detection signal with a lowenergy due to the scattering and a noise and reduce the load of signalprocessing of the data processing device 12 and prevent degradation ofimage quality. The present embodiment makes it also possible to find, atan early stage, the address of a abnormal detector 21, which can notdistinguish a γ-ray detection signal of a low energy from a noise, andthen automatically exclude the same.

Then, in this noise determination, an additional specific analog ASIC 24circuit is not required and the output signals of the slow systems 24B,24C for making the γ-ray detection time more accurate just need to beused. This is advantageous in terms of cost of the signal processingdevice.

In particular, in the conventional technique described inJP-A-2006-98411, an abnormality is determined based on a ratio of thecounting rates between detectors that are disposed in multilayer in theradial direction, while the abnormality determination in the presentembodiment can be carried out based on an output signal from onedetector 21. Accordingly, in the case where the radiation detectors arenot disposed in multilayer in the radial direction as described later,the present embodiment can be applied also to a PET apparatus using ascintillator and a photomultiplier.

Moreover, also when there is a channel of the detector 21 thatfrequently outputs a noise, the noise counting part 36 f detects thechannel of the detector 21 outputting more noise count signals than aspecified reference value, and outputs the abnormality determination tothe control part 36 g, whereby the control part 36 g prevents theaddress calculation part 36 a from carrying out address calculation tothe output signal from the relevant detector 21 or from controlling ADC25A, 25B. Thus, the signal processing load of the digital ASIC 26 andFPGA 31 can be reduced.

In this way, a noise signal can be prevented from being outputted to thedata processing device 12 as a packet data from the digital ASIC 26 forimage generation, so that PET images with an excellent image quality canbe generated.

Moreover, the control part 36 g checks the noise count status at apredetermined cycle with respect to the detector 21 that has beendetermined as abnormal, and if the frequency of occurrence of the noisecount satisfies a specified reference value, the detector 21 isrecovered for PET image generation thus enabling automatic recover alsofrom a temporary detector abnormality.

Furthermore, the control part 36 g outputs to the operator console 13the date-time information obtained when determined as abnormal and thedetector ID for the detector 21 that is determined as abnormal.Accordingly, by viewing this information on the operator console 13 sideduring maintenance, an operator can determine a need for inspection ofthe detector unit 2 and monitor a distribution of abnormal detectorsamong the whole detectors 21.

In addition, since the correlation data stored in the noisedetermination part 36 e and the reference value data for the abnormalitydetermination used by the noise counting part 36 f can be changed fromthe operator console 13, reconfiguration, which flexibly responds to achange with time of the output signal of the detector 21 and a change inthe waveform of an output signal due to changes in the installationenvironment of the PET apparatus 1, or the like, can be carried out.

In addition, in the present embodiment, if the noise counting part 36 fdetermines a detector as abnormal, the control part 36 g carries outcontrol to prevent the detector control part 36 from processing therelevant signal thereafter, but not limited thereto. For example, inprocessing a detected γ-ray energy value for the relevant detector 21,for the detection γ-ray signals that have not being determined as anoise, the control can be carried out so as to process only thosecorresponding to an energy window obtained by providing a specifiedenergy to 511 keV of an incident γ-ray energy.

Second Embodiment

Next, a nuclear medical diagnosis apparatus, which is another embodimentconcerning the present invention, will be described with reference toFIG. 17 to FIG. 20. The nuclear medical diagnosis apparatus of thepresent embodiment is a SPECT apparatus.

The same configuration as that of the first embodiment is given the samereference numeral to omit the duplicated description.

FIG. 17 is a perspective view showing the configuration of a SPECTapparatus, and FIG. 18 is a block diagram showing a connection relationbetween an analog ASIC and a digital ASIC in the SPECT apparatus. FIG.19 is a functional block diagram of the analog ASIC, and FIG. 20 is afunctional block diagram of the digital ASIC.

A SPECT apparatus 51 comprises a pair of radiation camera parts 52, arotating support stand 57, a data processing device 58, and an operatorconsole 13A. The radiation camera parts 52 are disposed facing to eachother on the rotating support stand 57 at positions shifted by 180° inthe circumferential direction. Specifically, each unit supporting member56 of the respective radiation camera parts 52 is mounted to therotating support stand 57 at a position separated by 180° in thecircumferential direction. A plurality of detector units 102 eachincluding twelve coupling substrates are removably mounted to therespective unit supporting members 56.

The detector 21 is held in the detector unit 102. The configuration ofthe respective detector units 102 is the same as the configuration ofthe detector unit 2 in the first embodiment except the configuration ofa coupling substrate 120.

The coupling substrate 120 includes a detector substrate 120A and anASIC substrate 120B as in the coupling substrate 20 of the firstembodiment (see FIG. 19). The detector 21 positioned at a tip portion ofeach detector substrate 120A is positioned on the bed 14 side. Acollimator 55 formed of a radiation shielding material, e.g., lead,tungsten, or the like, is provided on the test object P side of therespective radiation camera parts 52. Each collimator 55 forms a largenumber of radiation paths through which γ-rays pass. These radiationpaths are provided one-to-one corresponding to each detector 21positioned at the tip portion of each detector substrate 120A of oneradiation camera part 52. The coupling substrate 120 and the collimator55 are disposed inside a light shield/electromagnetic shield 54 disposedin the rotating support stand 57. The collimator 55 is mounted to thelight shield/electromagnetic shield 54. The light shield/electromagneticshield 54 blocks influence on the detector 21 and the like fromelectromagnetic waves other than γ-rays.

The bed 14, on which the test object P dosed with radiopharmaceuticallies, is moved, and the test object P is moved between a pair ofradiation camera parts 52. As the rotating support stand 56 is rotated,each radiation camera part 52 pivots around the test object P. A γ-rayemitted from an accumulation part inside the test object P, where theradiopharmaceutical gathers, e.g., an affected part, is incident uponthe corresponding detector 21 through a radiation path of the collimator55. When the γ-ray interacts with the detector 21, the detector 21 willoutput a γ-ray detection signal. This γ-ray detection signal isprocessed by the later-described analog ASIC 124 and digital ASIC 126.

Because the configuration of the detector substrate 120A used in thepresent embodiment is the same as the configuration in the firstembodiment, the description in the present embodiment is omitted. TheASIC substrate 120B constituting the coupling substrate 120 will bedescribed with reference to FIG. 18 to FIG. 20.

The ASIC substrate 120B is connected to the detector substrate 120A viathe connector C1 as in the coupling substrate 20 in the firstembodiment. The ASIC substrate 120B includes the capacitor 22 andresistor 23, four analog ASICs 124, and one digital ASIC 126 providedfor each detector 21.

The analog ASIC 124 comprises the preamplifier 24 a, the first slowsystem 24B, the second slow system 24C, and 32 sets of analog signalprocessing circuits 133 having a trigger output circuit 24 h. The analogsignal processing circuit 133 is provided for each detector 21.

Here, since the SPECT apparatus 51 does not carry out simultaneousmeasurement of a γ-ray pair, a fast γ-ray detection trigger signal isnot required. Accordingly, the output signal VS1 of the waveform shapercircuit 24 d of the first slow system 24B is used to input the outputsignal VS1 to the trigger output circuit 24 h. In order to removeinfluence of noise, the trigger output circuit 24 h outputs a triggersignal VT′ upon input of the γ-ray detection signal VS1 equal to orgreater than a setting level.

The digital ASIC 126 includes 16 sets of detection signal processingparts 134 and one data transfer part 37A, wherein each detection signalprocessing part 134 includes eight timing detection parts 135 and onedetector control part 136. The digital ASIC 126 is an LSI integratingthese. All the digital ASIC 126 provided in the SPECT apparatus 51receive a clock signal from a non-illustrated 64 MHz clock generator(quartz oscillator) to operate in a synchronous manner. The clock signalinputted to each digital ASIC 126 is inputted to the respective timingdetection part 135 and detector control part 136 inside all thedetection signal processing parts 134.

The timing detection part 135 is provided for each detector 21, and thetiming signal VT′ is inputted from the trigger output circuit 24 h ofthe corresponding analog signal processing circuit 133. The timingdetection part 135 determines a detection time of a γ-ray based on theclock signal when the timing signal VT′ is inputted, and generatesdetection time information.

The detector control part 136 includes the address calculation part 36a, the detection energy correcting part 36 c, a packet data generationpart 136 d, the noise determination part 36 e, the noise counting part36 f, and the control part 36 g.

Upon receipt of the detection time information corresponding to thetiming signal VT′ obtained when a γ-ray is detected, from the timingdetection part 135, the address calculation part 36 a identifies arelevant detector ID, and outputs the detector ID and the detection timeinformation to the detection energy correcting part 36 c, the packetdata generation part 136 d, and the noise determination part 36 e. Thatis, the address calculation part 36 a stores the detector IDcorresponding to each timing detection part 135 connected to the addresscalculation part 36 a, whereby when the detection time information isinputted from a certain timing detection part 135, the addresscalculation part 36 a can identify an detector ID corresponding to thetiming detection part 135. This is possible because the timing detectionpart 135 is provided for each detector 21.

Furthermore, after the trigger signal is inputted the addresscalculation part 36 a, the address calculation part 36 a outputs a peakhold control signal to the analog signal processing circuit 133including the above-described identified detector ID, and also outputsthe detector ID and the ADC control signal to ADCs 25A, 25B. Uponreceipt of the peak hold control signal, the peak hold circuits 24 e, 24g of the analog signal processing circuit 133 carry out peak holdprocessing to the signal inputted from the waveform shaper circuits 24d, 24 f. Then, upon receipt of a reset signal from the addresscalculation part 36 a after a predetermined time, the peak hold circuits24 e, 24 g cancel the peak hold processing. The ADCs 25A, 25B convertthe pulseheight values (voltage values) VE1, VE2 outputted from the peakhold circuits 24 e, 24 g of the analog signal processing circuit 133corresponding to the detector ID inputted from the address calculationpart 36 a into a digital signal and output the same to the noisedetermination part 36 e. These pulseheight values VE1, VE2 are inputtedto the noise determination part 36 e.

The noise determination part 36 e determines whether the relevantdetection signal is a noise signal or a γ-ray detection signal based onthe correlation between two pulseheight values VE1, VE2. If determinedas a γ-ray detection signal, the noise determination part 36 e willoutput the pulseheight values VE1, VE2 to the detection energycorrecting part 36 c but will not output a noise count signal to thenoise counting part 36 f.

By using a correction value corresponding to the gains and offsets ofthe detector 21 and analog ASIC 124 corresponding to the detector IDinputted from the address calculation part 36 a, the detection energycorrecting part 36 c calculates pulseheight values VE1′, VE2′ that arecorrected from the pulseheight values VE1, VE2, and generatesinformation on the detection energy value corresponding to the energy ofthe detection γ-ray, and outputs the same to the packet data generationpart 136 d.

When determined as a noise signal, the noise determination part 36 eoutputs a noise count signal to the noise counting part 36 f along withthe detector ID, and outputs a reset signal to the packet datageneration part 36 d.

(Operation of Detector Output Signal Processing Control Part)

The noise counting part 36 f and the control part 36 g have the samefunctions as those of the first embodiment and operate in the samemanner.

The packet data generation part 136 d appends the detector ID and thedetection time information to the information on the detection energyvalue to thereby generate as a packet data (information on the detectedγ-ray, information on detected radiation), and outputs the same to thedata transfer part 37A. The data transfer part 37A periodicallytransmits the packet data, which is digital information outputted fromthe packet data generation part 36 d of each detection signal processingpart 134, for example, to one FPGA 31A provided outside the enclosure ofthe detector unit 102 (see FIG. 17) that houses twelve couplingsubstrates 120 therein. FPGA 31A transmits the digital information tothe data processing device 58 via information transmission wiringconnected to the connector 38A.

A rotation angle detected by an angle gauge (not shown) connected to therotating shaft of a motor (not shown) that rotates the rotating supportstand 57 is inputted to the data processing device 58. This rotationangle indicates the rotation angle of each radiation camera part 52,specifically indicating the rotation angle of each detector 21. On thebasis of this rotation angle, the data processing device 58 calculatesthe position (positional information) on the pivoting track of eachpivoting detector 21. Accordingly, a position (position coordinate) ofthe detector 21 obtained when a γ-ray is detected can be calculated. Onthe basis of the detector ID that detected a γ-ray, the data processingdevice 58 counts a γ-ray for which the detection energy value is equalto or greater than a set value. The detection energy value here is asummation of the detection energy value of each γ-ray detection signalif there is coincidence in a plurality of detectors 21 (four detectors21 arranged side by side in a straight line in the (a) of FIG. 6)positioned on an extension of the radiation path of the collimator 55.This counting is carried out on each of the areas corresponding toincrements of 0.50° with respect to the rotational center of therotating support stand 57.

In addition, the data processing device 58 uses the positionalinformation on the detector 21 and the count value (count information)of γ-rays both obtained when the γ-ray was detected to prepare tomograminformation on the test object P for a position at which theradiopharmaceutical is concentrated, i.e., at the position of amalignant tumor. The tomogram information is displayed on the displaydevice 13 a. Information, such as the above-described packet data, countvalue and positional information on the detector 21 obtained through thesimultaneous measurement, and the tomogram information, is stored in astorage device of the data processing device 58.

Moreover, according to the present embodiment, when determined as anoise in the noise determination part 36 e, generating a packet data forthe relevant signal and sending this packet data to the data processingdevice 58 are not allowed to be carried out, and therefore thesignal-processing load of the digital ASIC 126 and FPGA 31 can bereduced.

Especially, in the conventional technique described in JP-A-2006-98411,an abnormality is determined based on a ratio of the counting ratesbetween detectors disposed in multilayer, while the abnormalitydetermination in the present embodiment can be made based on an outputsignal from one detector 21. Accordingly, in the case where theradiation detectors are not disposed in multilayer in the radialdirection as described later, the present embodiment can be applied alsoto a PET apparatus using a scintillator and a photomultiplier.

Also when there is a channel of the detector 21 that frequently outputsnoise, the noise counting part 36 f detects a channel of the detector 21outputting more noise count signals than a specified reference value,and outputs the abnormality determination to the control part 36 g,whereby the control part 36 g prevents the address calculation part 36 afrom carrying out address calculation to the output signal from therelevant detector 21 or from controlling ADC 25A, 25B. Accordingly, thesignal processing load of the digital ASIC 126 and FPGA 31A can bereduced.

In this way, a noise signal can be prevented from being outputted fromthe digital ASIC 126 to the data processing device 58 as a packet datafor image generation, so that SPECT images with an excellent imagequality can be generated.

Moreover, the control part 36 g checks the state of noise counts at apredetermined cycle for the detector 21 that has been determined asabnormal, and if the frequency of occurrence of a noise count satisfiesa specified reference value, the detector 21 is recovered for SPECTimage generation, thus enabling automatic recovery also from a temporarydetector abnormality.

Furthermore, the control part 36 g outputs to the operator console 13Athe date-time information obtained when determined as abnormal and thedetector ID for the detector 21 determined as abnormal. Accordingly, byviewing this information on the operator console 13A side duringmaintenance, an operator can determine a need for inspection of thedetector unit 102 and the unit substrate and monitor a distribution ofabnormal detectors 21.

In addition, since the correlation data stored in the noisedetermination part 36 e and the reference value data for abnormalitydetermination used by the noise counter 36 f can be changed from theoperator console 13A, reconfiguration that flexibly responds to a changein the waveform of an output signal due to a change with time of theoutput signal of the detector 21, a change in the installationenvironment of the SPECT apparatus 51, or the like, can be carried out.

In addition, in the first embodiment and the second embodiment, althoughpulseheight values outputted from the peak hold circuits are used, asampled pulseheight value based on a timing signal may be used instead.

Incidentally, in the first embodiment and the second embodiment, thefirst and second slow systems 24B, 24C used in noise determination havea configuration wherein the first slow system 24B is used formeasurement of an intended γ-ray detected energy and this first slowsystem 24B is also used for noise determination at the same time. Thisis for simplification of the circuit configuration. Two slow systemshaving mutually different time constants used for noise determinationmay be prepared totally independently of the slow system used fordetection energy measurement of an intended γ-ray, i.e., three slowsystems, may be prepared for one detector 21.

In addition, in the above-described first embodiment and secondembodiment, the PET apparatus 1 and the SPECT apparatus 51 using thesemiconductor radiation detector 21 as a radiation detector have beendescribed, but the present invention is not limited thereto.

The present invention may be applied also to a PET apparatus and SPECTapparatus using a γ-ray detector obtained by combining theabove-described scintillator of NaI or the like with a photomultiplieror a photodiode.

In this case, a γ-ray detector that employs a scintillator in the radialdirection with respect to the body axis of the test object P is notdisposed in multilayer but is dispose in one layer. Here, a plurality ofγ-ray detectors are usually disposed in the circumferential direction aswell as in the body axis direction.

In this case, a photomultiplier or a photodiode receives scintillationlight outputted from a scintillator and then converts this into anelectrical signal corresponding to the quantity of light. Accordingly, adifference in the detection signal waveforms resulting from a differencein the mobility between a hole and an electron will not occur unlike inthe case of semiconductor radiation detectors. Accordingly, for thepulseheight values VE1, VE2 outputted from the first slow system and thesecond slow system, respectively, having mutually different timeconstants, if the pulseheight values VE1, VE2 are included in a region84 shown in FIG. 21, the noise determination part 36 e determines theoutput signal from a γ-ray detector as an intended γ-ray detectionsignal, while if these are included in a region 80, the noisedetermination part 36 e determines this as a noise. Then, the noisecounting part 36 f and the control part 36 g can function based on thisdetermination result, as shown in the first embodiment and the secondembodiment.

Moreover, the method for determining whether the output signal from aγ-ray detector is an intended γ-ray detection signal or a noise is notlimited to the comparison of the pulseheight values of the outputsignals from the waveform shaper circuits having mutually different timeconstants described in the first embodiment and second embodiment. Forexample, as shown in FIG. 5 of the published literature “ASTRONOMY &ASTROPHYSICS SUPPLEMENT SERIES” (122, 357-369, (1997)), thediscrimination may be carried out based on a correlation between thepulseheight value of an output signal from a γ-ray detector and thewaveform selection (for example, based on the rising characteristics ofthe waveform).

In addition, in each of the above-described embodiments, the PETapparatus 1 and the SPECT apparatus 51 have been described, but thepresent invention may be applied to a γ camera, as well. The γ cameraprovides two-dimensional functional images and comprises a collimatorrestricting the incident angle of a γ-ray.

In addition, the PET apparatus 1 or the SPECT apparatus 51 may becombined with an X ray CT to configure a nuclear medical diagnosisapparatus.

In the nuclear medical diagnosis apparatus, a γ-ray from a test objectmay be scattered in a certain radiation detector and absorbed by anotherradiation detector, thereby providing energy to a plurality of radiationdetectors. In such a case, as disclosed in JP-A-2003-255048, it may bedetermined, on the basis of the radiation detection information in twoor more radiation detectors, whether a γ-ray before being scattered is aγ-ray from the radiopharmaceutical that is applied to a test object P,and if so, this γ-ray may be processed as an effective signal.Hereinafter, such a method is referred to as scattering radiationprocessing. The nuclear medical diagnosis apparatus includes ascattering radiation processing unit that identifies a plurality ofradiation signals, the plurality of radiation signals resulting fromradiation that scattered in a radiation detector, as one radiationsignal based on an output signal outputted from a noise determinationunit. Here, the scattering radiation processing unit may be any one ofthe packet data generation part 36 d, FPGA 31, and the data processingdevice 12. Since FPGA 31 has more radiation signals than the packet datageneration part 36 d, it carries out the scattering radiation processingto a wider target area, and since the data processing device 12 has moreradiation signals than FPGA 31, it carries out the scattering radiationprocessing to a wider target area. A noise determination unit on theupstream side of the scattering radiation processing unit may determinea relevant radiation signal as not being a noise, and then thescattering radiation processing unit may carry out scattering radiationprocessing based on an output signal after the determination. Thereby,the load of scattering radiation processing can be reduced. Moreover,because the scattering radiation processing increases effective signals,an accurate diagnostic image is expected to be obtained.

In the scattering radiation processing, energy information is one of thecritical information in determining whether a γ-ray before scattering isa γ-ray coming from the radiopharmaceutical that is applied to the testobject P. For this reason, a semiconductor detector excellent in energyresolution is preferably used as the radiation detector.

In carrying out the scattering radiation processing, an energy detectedin a radiation detector is smaller than that of the original γ-ray.Generally, the noise signals emitted by a faulty detector mostly have alow energy, so even a noise signal, which is not problematic when thescattering radiation processing is not carried out, may result in acritical failure in carrying out the scattering radiation processing.According to the present invention, the contribution of a noise signalcan be suppressed regardless of the magnitude of the energy, so that anaccurate diagnostic image can be provided.

It should be further understood by those skilled in the art thatalthough the foregoing description has been made on embodiments of theinvention, the invention is not limited thereto and various changes andmodifications may be made without departing from the spirit of theinvention and the scope of the appended claims.

1. A nuclear medical diagnosis apparatus, comprising an image pickupdevice including: a plurality of radiation detectors that detectradiation from a test object or radiation passing through the testobject; and a signal processing device connected to each of theradiation detectors, the signal processing device processing an outputsignal from the radiation detector, the nuclear medical diagnosisapparatus generating an image based on information on detected radiationoutputted from the signal processing device of the image pickup device,wherein the signal processing device comprises: a determination unitthat determines whether an output signal from the radiation detector isan intended radiation detection signal or a noise; a counting unit thatcounts for the radiation detector a number of times the output signalfrom the radiation detector is determined as a noise; and a control unitwhich, based on the number of times the output signal from the detectoris determined as a noise, determines a relevant radiation detector asfaulty, and which controls so as not to process an output signal fromthe radiation detector that is determined as faulty.
 2. The nuclearmedical diagnosis apparatus according to claim 1, wherein the signalprocessing device includes two waveform processing circuits havingmutually different time constants that carry out waveform processing ofan output signal from the radiation detector, and wherein thedetermination unit determines based on output signals from the twowaveform processing circuits.
 3. The nuclear medical diagnosis apparatusaccording to claim 1, wherein the control unit controls so as not toprocess an output signal from the relevant radiation detector by cuttingoff a power supply supplied to the radiation detector that is determinedas faulty.
 4. The nuclear medical diagnosis apparatus according to claim1, wherein the control unit appends an identification data to theinformation on detected radiation, the identification data indicative ofa faulty signal, for the radiation detector that is determined asfaulty, and wherein by making identifiable in generating the image,control is carried out so as not to process an output signal from therelevant radiation detector.
 5. A nuclear medical diagnosis apparatus,comprising: a supporting member in which a plurality of opening portionsare formed in a circumferential direction; and a plurality of detectorunits each being inserted in the opening portions, respectively, to beremovably mounted to the supporting member, wherein the detector unitincludes a housing member having light shielding characteristics, and aplurality of unit substrates removably housed in the housing member,wherein the unit substrate includes a plurality of semiconductorradiation detectors upon which a γ-ray is incident, and an integratedcircuit that processes a γ-ray detection signal outputted by each of theplurality of semiconductor radiation detectors, wherein the integratedcircuit outputs information on the detected γ-ray containing addressinformation on the semiconductor radiation detector that detects aγ-ray, wherein the detector unit includes another integrated circuitthat outputs to the outside of the detector unit the information on thedetected γ-ray outputted from the integrated circuit contained in eachof the unit substrates, the other integrated circuit being disposed inan external surface of the housing member, wherein the integratedcircuit that processes the γ-ray detection signal comprises: adetermination unit that determines whether an output signal from thesemiconductor radiation detector is an intended γ-ray detection signalor a noise; a counting unit that counts the number of times the outputsignal from the detector is determined as a noise, for each of thesemiconductor radiation detectors; and a control unit which, based onthe number of times the output signal from the detector is determined asa noise, determines a relevant radiation detector as faulty, and whichcontrols so as not to process an output signal from the radiationdetector that is determined as faulty.
 6. The nuclear medical diagnosisapparatus according to claim 5, wherein the integrated circuit forprocessing the γ-ray detection signal includes two waveform processingcircuits having mutually different time constants that carry outwaveform processing to an output signal from the radiation detector, andwherein the determination unit determines based on output signals fromthe two waveform processing circuits.
 7. The nuclear medical diagnosisapparatus according to claim 5, wherein the control unit controls so asnot to process an output signal from a relevant semiconductor radiationdetector by cutting off a power supply supplied to the semiconductorradiation detector that is determined as faulty.
 8. The nuclear medicaldiagnosis apparatus according to claim 5, wherein the control unitappends an identification data to the information on the detected γ-ray,the identification data indicative of a faulty signal, for thesemiconductor radiation detector that is determined as faulty, andwherein by making identifiable in generating the image, control iscarried out so as not to process an output signal from the relevantradiation detector.
 9. A nuclear medical diagnosis apparatus, comprisingan image pickup device including: a plurality of radiation detectorsthat detect radiation from a test object or radiation passing throughthe test object; and a signal processing device connected to each of theradiation detectors, the signal processing device processing an outputsignal from the radiation detector, the nuclear medical diagnosisapparatus generating an image based on information on detected radiationoutputted from the signal processing device of a relevant image pickupdevice, wherein the signal processing device includes a noisedetermination unit which determines whether an output signal from theradiation detector is an intended radiation detection signal or a noise,and which controls so as not to process the output signal if determinedas a noise.
 10. The nuclear medical diagnosis apparatus according toclaim 9, comprising: a supporting member in which a plurality of openingportions are formed in a circumferential direction; and a plurality ofdetector units each being inserted in the opening portions,respectively, to be removably mounted to the supporting member, whereinthe detector unit includes a housing member having light shieldingcharacteristics, and a plurality of unit substrates removably housed inthe housing member, wherein the unit substrate includes the plurality ofradiation detectors, and an integrated circuit that processes a γ-raydetection signal outputted by each of the plurality of radiationdetectors, wherein the integrated circuit outputs information on thedetected γ-ray containing address information on the radiation detectorthat detects a γ-ray, wherein the detector unit includes anotherintegrated circuit that outputs to the outside of the detector unit theinformation on the detected γ-ray outputted from the integrated circuitcontained in each of the unit substrates, the other integrated circuitbeing disposed in an external surface of the housing member, and whereinthe integrated circuit that processes the γ-ray detection signalincludes the noise determination unit.
 11. The nuclear medical diagnosisapparatus according to claim 9, wherein the signal processing deviceincludes two waveform processing circuits having mutually different timeconstants that carry out waveform processing to an output signal fromthe radiation detector, and wherein the determination unit determinesbased on output signals from the two waveform processing circuits. 12.The nuclear medical diagnosis apparatus according to claim 9, wherein atleast one of a determination condition for a determination whether theintended radiation detection signal or a noise in the determinationunit, and a condition for a determination as faulty, can be set by aninput unit from the outside.
 13. The nuclear medical diagnosis apparatusaccording to claim 9, capable of displayably outputting to an operatorfault information on a radiation detector that is determined as faulty.14. The nuclear medical diagnosis apparatus according to claim 9,wherein the radiation detector is a semiconductor radiation detectorusing a semiconductor element.
 15. The nuclear medical diagnosisapparatus according to claim 1, wherein at least one of a determinationcondition when the determination unit determines whether an outputsignal from the radiation detector is the intended radiation detectionsignal or a noise, and a condition when the determination unitdetermines a relevant radiation detector as faulty, can be set by aninput unit from the outside.
 16. The nuclear medical diagnosis apparatusaccording to claim 5, wherein at least one of a determination conditionwhen the determination unit determines whether an output signal from theradiation detector is the intended radiation detection signal or anoise, and a condition when the determination unit determines a relevantsemiconductor radiation detector as faulty, can be set by an input unitfrom the outside.
 17. The nuclear medical diagnosis apparatus accordingto claim 1, capable of displayably outputting to an operator faultinformation on a radiation detector that is determined as faulty. 18.The nuclear medical diagnosis apparatus according to claim 5, capable ofdisplayably outputting to an operator fault information on a radiationdetector that is determined as faulty.
 19. The nuclear medical diagnosisapparatus according to claim 5, further comprising a tomograminformation preparation device that prepares tomogram information basedon the information on the detected γ-ray outputted from the integratedcircuit.
 20. The nuclear medical diagnosis apparatus according to claim9, further comprising a scattering radiation processing device which,based on an output signal outputted from the noise determination unit,identifies a plurality of radiation signals as one radiation signal, theradiation signals resulting from radiation that scatters in a radiationdetector.